1. Field of the Invention
The present invention concerns a method for determining a compensation setting for an eddy current field.
2. Description of the Prior Art
Magnetic resonance technology is a known modality for acquiring images of the inside of a body of an examination subject. In a magnetic resonance device, rapidly switched gradient fields that are generated by a gradient coil system are superimposed on a static homogenous basic magnetic field that is generated by a basic field magnet. The magnetic resonance device also has a radio-frequency system that radiates radio-frequency signals into the examination subject to excite magnetic resonance signals and that acquires the generated magnetic resonance signals, on the basis of which magnetic resonance images are created.
A gradient coil of the gradient coil system generates a gradient field for a specific direction in space that, in the desirable ideal case, has (at least within an imaging volume) a field component that is collinear to the basic magnetic field. The field component has a predeterminable gradient that is spatially-independently, optimally of same magnitude at an arbitrary particular point in time, at least within the imaging volume. Since the gradient field is a temporally variable magnetic field, the aforementioned is true for each individual point in time, but the strength of the gradient is variable from one point in time to another point in time. To generate the gradient field, an appropriate current is adjusted (set) in the gradient coil. The amplitudes of the required currents amount to more than 100 A. The current rise and fall rates amount to more than 100 kA/s. For power supply, each gradient coil is connected to a gradient amplifiers.
The gradient coil system normally is surrounded by a conductive structure, and thus eddy currents are induced in this structure by the switched gradient fields. Examples of such conductive structures are a vacuum vessel and/or a cryoshield of a superconducting basic field magnet, a copper foil of a radio-frequency shielding, and the gradient coil system itself. The eddy current fields generated by the eddy currents are undesirable because without counteracting measures, they weaken the gradient fields and distort their time curves. This leads to an impairment of the quality of magnetic resonance images. This is also true for an actively shielded gradient coil system that has shielding coils associated with the gradient coils, whereby a quantitative reduction of eddy currents is achieved in comparison to the unshielded gradient coil systems.
The impairment of a gradient field as a result of the eddy current fields can be compensated to a certain degree by a corresponding pre-distortion of a quantity used to control the gradient field. To compensate, the control quantity is filtered such that eddy current fields ensuing given non-predistorted operation of the gradient coil are cancelled by the pre-distortion. The eddy current fields can be described in the form of a series expansion of a spherical function. To describe the temporal dependency of the eddy current field components, a temporally falling exponential function characterized by a time constant is associated with each coefficient of the spherical function in the series expansion. For filtering, a suitable filter can be used with parameters determined by the time constants and coefficients.
Filters employed as high-pass filters are known from European Application 0 228 056. To determine the parameters necessary for the filter, the eddy current field is first measured. A method is specified for this purpose wherein the magnetic field curve is measured by the magnetic resonance signals induced in a probe. Since measurement of the eddy current field is necessary at at least two locations of the imaging volume, the probe must be switched back and forth between two measurement positions for each measurement cycle. Since, in magnetic resonance devices, it is furthermore necessary in many cases to record the existing eddy current field in as large an area of the imaging volume as possible, the measurement with the probe is therefore extremely complicated, since the entire imaging volume must be scanned for complete recording of the eddy current field. This is particularly true when it is also desired to measure eddy current field portions of a higher order. An automated measurement is not possible using this known technique.
A method is described in U.S. Pat. No. 6,025,715 in which the eddy current compensation is begun with various sets of filter parameters, and then the remaining (residual) fields are respectively measured. An optimal compensation is determined via empirical interpolation of the dependency of the remaining fields on the filter parameters. A disadvantage of this method is that 53 measurements per gradient field are necessary, and thus the method takes a very long time.
A method to measure eddy current fields is known from German OS 43 13 392 in which a volume-occupying phantom (as it is also used for other test and setting purposes in magnetic resonance devices) is introduced into the examination space and data are acquired with a slice-selective magnetic resonance method. An eddy current compensation can thereby be implemented and checked fully automatically without special devices such as measurement probes. The operation of the method is simple since the phantom does not have to be moved for the measurement.
In an development of the method of German OS 43 13 392, a method based on magnetic resonance measurements is described in German PS 198 59 501 with which terms known as cross-terms additionally can be determined. A cross term is a field component of the eddy current field that is caused by the gradient field with a gradient in a first direction, the field component acting in a second direction that is perpendicular to the first. If the field component is a field component of the first order, the field component can be compensated by a counter-directed triggering of a gradient coil, with which a gradient field in the second direction can be generated. In this method, a volume-occupying phantom is brought into the imaging volume of the magnetic resonance device, a measurement gradient pulse of predeterminable pulse width is switched, and after the deactivation of the measurement gradient pulse at least two imaging sequence blocks, temporally separate from one another, are generated, from each of which imaging magnetic resonance signals are acquired from these signals a (at least) two-dimensional data set is generated, with the phase information contained in the magnetic resonance signals embodying characteristics (parameters) of the eddy current fields. The amplitudes and time constants of the eddy current field can be determined from these characteristics with a suitable evaluation method.
Furthermore, in magnetic resonance devices the use of shim coils is known, with which the basic magnetic field can be homogenized, for example dependent on different examination subjects. For this purpose, the shim coils are operated with suitable direct currents. Since linear basic magnetic field deviations (meaning interferences of the first order) can be compensated (by the gradient coils being charged with a direct current), the shim coils normally are fashioned such that precisely one interference of a specific order larger than the first order is compensated with one of the shim coils with the respective shim coils compensating for interferences of different orders greater than one. Moreover, it is known from previously cited German PS 198 59 501 that eddy current-caused interferences of a higher order additionally can be compensated by an additional pulsed current feed of the shim pulses.
The eddy currents and the eddy current compensation can be mathematically described as follows with reference to FIG. 1: a desired gradient pulse u(t) as an input signal (in the simplest case, a rectangular pulse as shown in FIG. 1) is predistorted by a eddy current compensation filter. The transfer function of the filter is K(t). The gradient amplifier generates a current proportional thereto. Assuming the amplification does not distort the signal shape, it does not have to be considered further in the mathematical description. The gradient coil now generates a gradient field g(t) as an output signal. Eddy currents in the surrounding electrically-conductive structures act in a counteracting manner and distort the pulse shape. The effect of the eddy currents can be described by a transfer function W(t). Given perfect eddy current compensation, the gradient field g(t) has the same pulse shape as the input signal u(t). by convolution of the input signal u(t) with the transfer functions K(t) and W(t), the gradient field g(t) mathematically yields:g(t)=u(t)*K(t)*W(t)After a Laplace transformation, the convolution operation becomes a simple multiplication:{tilde over (g)}(s)=ũ(s){tilde over (K)}(s){tilde over (W)}(s)The variable of the Laplace transformation is designated s, and the Laplace-transformed quantities are designated with “{tilde over ( )}” over the letters, with use of the same letters. For a perfect eddy current compensation, the output signal g(t) should correspond to the input signal u(t). The transfer function of the eddy current compensation filter K(t) thus results from the following equation:             K      ~        ⁡          (      s      )        =      1                  W        ~            ⁡              (        s        )            According to the known method, the eddy current compensation ensues as follows: the eddy current compensation is initially disabled, thus {tilde over (K)}(s)=1. A stepped input signal u(t) is applied, for which the Laplace-transformed ũ(s)=1/s. An eddy current field g(t) that can be described as a sum of exponential functions arises as a reaction to this stepped input signal. The Laplace-transform {tilde over (g)}(s) can be analytically or numerically calculated from this. The eddy current transfer function thus results as:                     W        ~            ⁡              (        s        )              =                                        g            ~                    ⁡                      (            s            )                                                u            ~                    ⁡                      (            s            )                              =              s        ⁢                              g            ~                    ⁡                      (            s            )                                ,and the desired transfer function of the eddy current compensation filter:             K      ~        ⁡          (      s      )        =            1                        W          ~                ⁡                  (          s          )                      =          1              s        ⁢                              g            ~                    ⁡                      (            s            )                              The filter parameters can be calculated from the filter transfer function.
Finally, a magnetic resonance device with a gradient coil system is known from German OS 101 56 770 in which an electrically-conductive structure is arranged and fashioned such that, at least within the imaging volume of the magnetic resonance device, a magnetic field of the structure caused by a gradient field due to induction effects is similar (in the geometric sense) to the gradient field. In an embodiment, at least one part of the structure is a barrel-shaped component of the basic field magnet. Among other things, an advantage achieved with this arrangement is that the gradient coil system can be fashioned without secondary coils, since the undesirable consequences of the switched gradient fields due to the geometric similarity of the magnetic field caused by the structure can be controlled by a pre-distortion.